Method and system for systemic delivery of growth arresting, lipid-derived bioactive compounds

ABSTRACT

A system and method for optimizing the systemic delivery of growth-arresting lipid-derived bioactive drugs or gene therapy agents to an animal or human in need of such agents utilizing nanoscale assembly systems, such as liposomes, resorbable and non-aggregating nanoparticle dispersions, metal or semiconductor nanoparticles, or polymeric materials such as dendrimers or hydrogels, each of which exhibit improved lipid solubility, cell permeability, an increased circulation half life and pharmacokinetic profile with improved tumor or vascular targeting.

The present application claims priority to U.S. Provisional ApplicationNo. 60/465,938, filed Apr. 25, 2003, which is incorporated herein byreference.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to the field of nanotechnology. Moreparticularly, the present invention provides nanoscale assembly systemsfor systemic delivery of therapeutic bioactive lipid compounds and/orhydrophobic chemotherapeutic agents and/or nucleotide/gene agents toindividuals in need of such therapy.

2. Description of Related Art

Nanotechnology has been intricately linked with the life sciences(generally referred to as nanobiotechnology) since its inception byRichard Keynman in his 1959 speech, “There's Plenty of Room at theBottom,” in which he made reference to the complexity and smallness ofthe living cell and challenged the scientific community to “make a thingvery small which does what we want” (Feynman, R. P., 1959, available:www.zyvex.com/nanotech/feynman.html). Although commercialnanobiotechnology is still in its infancy, the rate of nanoscaleassembly system development has been increasing exponentially in thelast ten years, due to the unique advantages that these systems offerfor drug delivery and therapeutics. Examples of some nanoscale assemblysystems include liposomes, polymeric structures such as dendrimers andhydrogels, and metal or semiconductor nanoparticles referred to asquantum dots.

Many effective reagents are available for introducing transcriptionallyactive DNA, and even functional peptides and proteins into viable cells.However, approaches to deliver bioactive lipids into living cells arenot generally available. The delivery of bioactive sphingolipids andphospholipid metabolites, analogues, mimetics or derivatives and theirintercalation into cells is impeded by their physical-chemicalproperties that render these lipids hydrophobic and cell impermeable.

Ceramide, a sphingolipid that acts as a lipid-derived second messengerthat modulates the induction of cell differentiation, cell cycle arrestand/or apoptosis, is an example of a bioactive lipid whose exogenousadministration has been problematic. Intracellular ceramide accumulationresults from multiple stimuli, such as growth factor deprivation,cytokines, chemotherapy and other cytotoxic agents, ionizing radiation,heat shock, and various environmental factors. These stimuli have beenobserved to initiate ceramide-mediated signaling cascades, including theinhibition of Akt pro-survival pathways and the stimulation of caspaseactivity, which ultimately leads to DNA fragmentation and cell death.Thus, based on ceramide's potent regulation of cell growth,differentiation, and death, and the fact that it is a natural moleculethat targets discrete kinases and signaling pathways linked toproliferation and/or survival, ceramide has been identified as atherapeutic agent in cancer and cardiovascular disease.

The clinical utility of local delivery of a cell-permeable ceramideanalogue, C₆, from drug-eluting platforms previously has beendemonstrated by Charles et al. (Circ. Res. 2000 Aug. 18:87(4):282-8).Specifically, ceramide-coated balloon catheters were shown to inducecell cycle arrest in stretch-injured vascular smooth muscle cells.Although the delivery of C₆-ceramide from coated and distended balloonsallow for direct delivery to the vasculature, there are severalobstacles to the delivery of ceramide for systemic applications, such ascancer chemotherapy or targeting diffuse atherosclerotic lesions andvulnerable plaque. In particular, three significant barriers to systemicceramide delivery exist, despite the use of short chain, more cellpermeable derivatives.

First, short-chain, cell-permeable ceramide analogues such as C₂, C₆,and C₈-ceramide are still lipids, and thus extremely hydrophobic bynature, precipitating as fine lipid micelle suspensions when added, inDMSO or ethanol vehicle, to cell media. Second, although short-chainceramide analogues are more cell-permeable than long-chain physiologicalceramide (C₁₈-C₂₄-ceramide), their sphingoid backbone limits theirintercalation into plasma membranes. Finally, the existence ofcirculating and intracellular ceramidases promote the conversion ofbioactive ceramides into less pro-apoptotic metabolites.

Organic solvent systems have been investigated in order to augment thedelivery of ceramide to cells. It has been proposed that adodecane/ethanol solvent system, which is insoluble in culture media,precipitates out with the ceramide and forms very small droplets, ormicelles, that fuse with the plasma membrane. The use of suchprecipitating solvents may be limited by the variability in particlesize and access to cellular membranes. Protein adjuvants, such as bovineserum albumin, may also assist in vitro ceramide delivery vianon-specific lipid/protein interactions, but would not permit theefficient delivery of sufficient quantities of C₆-ceramide to systemictargets.

Thus, in order to realize the therapeutic benefits of bioactive lipidsor gene therapy agents, there exists a need for improved systemicdelivery systems of such hydrophobic or charged chemotherapeuticcompounds into living cells of animals or humans in need of suchtherapy.

SUMMARY OF THE INVENTION

The present invention addresses this critical need by providing a systemand method for optimizing the systemic delivery of growth-arresting,pro-apoptotic, lipid-derived bioactive drugs and/or chemotherapeutichydrophobic agents and/or gene therapy agents to an animal or human inneed of such agents utilizing nanoscale assembly systems.

The present invention provides a method and system for maximizing thesystemic delivery of growth-arresting, pro-apoptotic, lipid-derivedbioactive therapeutic compounds and/or gene therapy agents to livingcells of an animal or human in need of such therapy, utilizing nanoscaleassembly systems, such as liposomes, resorbable and non-aggregatingdispersed nanoparticles, metal or semiconductor nanoparticles orpolymeric materials such as dendrimers or hydrogels, each of whichexhibit improved lipid solubility, cell permeability, an increasedcirculation half life and pharmacokinetic profile with improved tumor orvascular targeting.

In one embodiment of the present invention, polyethyleneglycol 450liposomes suitable for delivery of bioactive lipids, proteins andtherapeutic agents, referred to herein as “pegylated” liposomes areformulated that have one or more membranes comprised of agrowth-arresting lipid-derived bioactive compound and/or a gene therapyagent and/or cholesterol. These pegylated liposomes have been formulatedto contain PEG C8 (pegylated cell-permeable ceramide), ranging in sizebetween 750-5000 MW and/or PEG DSPE (disteroylphosphatidylethanolamine)ranging in size between 2000-5000 MW. The present embodiment uses PEG C8to stabilize the lipid bilayer, allowing the liposome to contain highmolar ration (i.e., 30%) of free bioactive C6 ceramide. In addition, theembodiment utilizes the PEG C8 as an integral component of the liposomethat contains the bioactive ceramide and/or a hydrophobicchemotherapeutic agent and/or a gene therapy agent. Moreover, PEG-C8formulated liposomes ensures optimal intercalation and localization ofthe free ceramide into caveolin-rich lipid rafts, a prerequisite formembrane internalization and transfer to subcellular organellesincluding the mitochondria for subsequent induction of apoptosis orprogrammed cell death of the targeted tissue or tumor. The pegylatedliposomes, also known as “stealth” liposomes, are capable of evadingclearance from the circulation by the reticuloendothelial system (RES),leading to improved circulation half life and tissue targeting.Targeting can be further achieved via the conjugation of particulartargeting moieties, such as antibodies and/or receptor ligands, whichwill promote the targeted accumulation into specific cells or tissues ofthe body. Additional embodiments assert that lipid therapeutics can alsobe formulated into “cationic” liposomes comprised of cationic lipids, inthe presence or absence of PEG-C₈, for effectively delivering negativelycharged oligonucleotides; or as “fusogenic” liposomes, in the presenceor absence of PEG-C₈, where the entire membrane of the liposome fuseswith the cell membrane of the target site to deliver the constituentsand contents of the liposome therein.

In another embodiment of the present invention, resorbable nanoparticleshaving a calcium phospho-silicate (CPS) shell are provided, in whichgrowth arresting, proapoptotic, lipid-derived bioactive compounds,and/or chemotherapeutic hydrophobic agents, and/or gene therapy agentsare loaded into the resorbable nanoparticles. The resorbablenanoparticles of the present invention can deliver chemotherapeutichydrophobic lipids or drugs or gene therapeutic agents systemically toliving cells, which normally are not transportable through thecirculation. A key feature of the synthesis of the resorbablenanoparticles is the proper dispersion (non-aggregation) of thenanoparticles in an aqueous liquid medium. One way to achieve dispersionis the use of size exclusion high performance liquid chromatography(SEC) modified specifically for the silicate-containing shellnanoparticles. Another way to achieve dispersion of the nanoparticles isto attach organic, inorganic or metal-organic dispersants to the outerCPS shell. Additionally, a carbodiimide-mediated polyethylene glycol(PEG) coupling agent can be attached to the alkylamine silane oralkylcarboxylic acid coupling agent to further ensure the “dispersed”non-aggregating state of the nanoparticles in vivo and to provide anconjugation point for targeting moieties onto the PEG coupling agent,thus enabling the nanoparticles to target specific sites forintracellular drug delivery.

In a further embodiment of the present invention, individual polymerscan be combined to form materials which are both “bio-smart”, i.e.respond to physical or chemical stimuli, and biodegradable in vivo andwhich can be loaded with the growth-arresting, lipid-derived bioactivecompounds and/or gene therapy agents.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1. Schematic of the preparation of the core-shell particles with aresorbable coating for drug delivery.

FIGS. 2A-B. Characterization of liposomal formulations. Liposomalformulations are produced with a spherical morphology and a homogeneoussize distribution. (A) Representative TEM of pegylated liposomalvesicles [DOPC/DOPE/CH/PEG-C₈/C6 (4:3:1:1:1)]. Identical micrographswere observed with conventional liposomal formulations [EPC/DOPE/CH/C₆(6:0.5:1.5:2)] (data not shown). Vesicular size was between 85 and 140nm in diameter; bar represents 100 nm. Extrusion of lipid solutions doesnot significantly diminish C₆ incorporation into liposomal vesicles. (B)Illustration of the average size of the liposomal formulations.

FIGS. 2C-D. Characterization of liposomal formulations. (C) Micelleformulations in a final concentration of 10 mg/ml, composed ofEPC/DOPE/CH/C6 (6:0.5:1.5:2) along with trace amounts of [³H]C₆ thatwere subjected to extrusion to produce conventional liposomes asdescribed. Mean±S.E., n=3 separate experiments. Lipid compositionremains consistent following extrusion of lipid micelle solution toproduce liposomal vesicles. (D) Representative TLC of conventionalliposomal formulations [EPC/DOPE/CH/C6 (3.5:3:2:1.5) andEPC/DOPE/CH/DHC₆ (3.5:3:2:1.5)] separated using a CHCL₃/MeOH/ddH₂O(60:25:5) solvent system. As expected, C₆ but not DHC₆ was stained withiodine due to the lack of the C₄₋₅ double bond of DHC₆.

FIG. 3A. In vitro pharmacokinetics of C₆ delivery. Liposomal delivery ofC₆ resulted in a greater cellular accumulation of C₆ as a function oftime than nonliposomal delivery. (A) Liposomes were formulated withtrace amounts of [³H]C₆ to determine the kinetics of ceramide deliveryto MDA cells. The total counts of liposomal and nonliposomal C₆ added tothe cells was set at 100%. At 20 μM, C₆ accumulation peaks atapproximately 16 h. Mean±S.E., n=3 separate experiments. *, p<0.05 whencomparing liposomal C₆ accumulation with nonliposomal C₆ accumulation.

FIGS. 3B-C. In vitro pharmacokinetics of C₆ delivery. Illustration thatliposomal C₆, but not Cholesteryl-1,2⁻³H(N) hexadecyl ether (3H-CHE)partitions into MDA cell membranes as a function of time (B) and dose(C). Pegylated liposomes [DOPC/DOPE/CH/PEG-C8/C₆ (4:3:1:1:1)] wereformulated with trace amounts of [³H C₆] and [³H CHE] to evaluate themechanism of ceramide (10 μM) delivery to MDA cells at the indicatedtime periods. A dose-dependent mechanism of ceramide delivery wasexamined over a 10-h treatment period. The mass of lipid delivered tocells was calculated as pmol/10⁶ cells. Mean±S.E., n=3 separateexperiments. *, p<0.05 when comparing C₆ accumulation with CHEaccumulation in respective formulations.

FIGS. 4A-B. Thymidine incorporation growth assays showing that liposomalC₆ delivery is more potent than non-liposomal C₆ in estrogenreceptor-negative MDA breast cancer cells. (A) conventional liposomes;(B) cationic liposomes.

FIG. 4C. Thymidine incorporation growth assay showing that pegylatedliposome C₆ delivery is more potent than non-liposomal C₆ in estrogenreceptor-negative MDA breast cancer cells.

FIG. 5. Pegylated liposomal C6[DSPC/DOPE/DSPC-PEG(5000)/C₈-PEG(750)/C₆-Cer (3.75:1.75:0.75:0.75:3.0)]delivery enhances the anti-proliferative activity of C6. Liposomaldelivery lowers the IC50 of C6 in 410.4 adenocarcinoma cells. Theincorporation of PEG-C₈ to 0.75 allows for the incorporation of 30 molepercent C₆. Results represent the mean±S.E. of three separateexperiments. *p<0.05.

FIG. 6. As a measure of apoptosis, liposomal C6 delivery augmentscaspase-3/7 activity in MDA cells.

FIG. 7A. Liposomal C₆ delivery augments the proapoptic activity ofintracellular C₆. TUNEL staining of fragmented 3′-OH DNA confirms thatC₆ treatment (20 μM) induces apoptosis in MDA cells. Apoptosis wasobserved to occur at approximately 16 h of incubation. Nonliposomal (20μM) C₆ and conventional liposomal C₆ [EPC/DOPE/CH/C₆ (6:0.5:1.5:2)] (20μM) induced DNA fragmentation in a similar manner to the Dnase-positivecontrol. Liposomal C₆ delivery results in a significant induction ofcellular apoptosis as measured by annexin V staining.

FIG. 7B. MDA cells were treated with nonliposomal C₆ (25 μM), pegylatedliposomal C₆ [DOPC/DOPE/CH/PEG-C8/C₆ (4:3:1:1:1)] (25 μM), or Ghostliposome for 24 h, stained with FITC-annexin V, and analyzed by flowcytometry. Mean±S.E., n=3 separate experiments. *, p<0.05; # p<0.005when compared with untreated control.

FIGS. 8A-B. Liposomal C₆ delivery modulates signaling cascadesassociated with growth inhibition and/or apoptosis. Liposomal C₆delivery inhibits Akt phosphorylation in MDA cells. (A and B) Cells werepretreated with nonliposomal C₆ (50 μM), pegylated liposomal C₆[DOPC/DOPE/CH/PEG-C8/C₆ (4:3:1:1:1)] (50 μM), or Ghost liposome for 8 hand then stimulated with 20 ng/ml IGF-1 for an additional 15 min.Protein lysates were probed for both native and active (phosphorylated)forms of Akt. (A) Representative blot of n=3 separate experiments. (B),mean±S.E., n=3 separate experiments. *, p<0.05 when compared withuntreated IGF-stimulated control.

FIGS. 9A-B. Liposomal C6 delivery results in the accumulation of C6 intocaveolae and mitochondrial structures. (A) Confocal microscopic imagesof NBD-C6 delivery to cells from liposomal vesicles demonstrates that C6also accumulates into cellular mitochondria. NBD-C6 co-localized withmitochondria. (B) Using [³H]-C6 as a marker for total C6, pegylatedliposomal delivery results in a time-dependent accumulation of ceramidein caveolae lipid signaling rafts. Ceramide accumulated in fractions#4-5 of a sucrose gradient, which represent caveolin-1 enriched lipidrafts (caveolae).

FIGS. 10A-B. Effect of pegylated liposomal C₆[DSPC/DOPE/DSPC-PEG(5000)/C8-PEG(750)/C₆-Cer (3.75:1.75:0.75:0.75:3.0)]on tumor volume. (A) The tumor volume of animals inoculated with 410.4adenocarcinoma cells was determined during and after treatment with 12,24 and 36 mg/kg liposomal C₆ and empty liposomal vehicles. Resultsrepresent the mean±S.E. of five animals per group. (B) Staining of tumorcryosections of tumor treatment for 1-week at 40 mg/kg, demonstratepositive TUNEL staining for apoptosis. Little staining is evident forGhost and untreated tumor sections. Representative slide from threeanimals per group and 10 random fields per tumor section.

FIG. 11A-B. Pharmacokinetics of pegylated liposomal C6 in 410.4 tumorbearing Balb/C mice. (A) 10 and 40 mg/kg doses of liposomal-C6 appear tofollow first order kinetics, with a sufficient plasma concentrationcorrelating to the in vitro IC50 sustained at 24 hours. (B) At thesedoses, a steady-state concentration of C6 in the tumor tissue isachieved at approximately 30 min. The 40 mg/kg dose maintains aconcentration well above the desired IC50 up to 24 hours.

FIG. 12. Proprietary dendritic structure composed of PLL, PLLA, andNIPAAM polymers that have thermo-responsive and biodegradableproperties.

FIG. 13A-B. Thermoresponsive and drug release properties of thedendrimers. A) Uvvis spectroscopy was used to study the transmittance ofsynthesized dendrimers at 0.5 and 0.1 mg/ml. A sharp transition insolution turbidity was observed at approximately 34° C., representingLCST of the dendrimers. B) C₆-loaded dendrimers display defined releasekinetics and in vitro bioefficacy. The fractional release of C₆ from theC₆-loaded dendrimers in distilled water containing 0.5% (w/v) SDS at 37°C. and 25° C. as a function of time. At 37° C., a temperature above theLCST of the dendrimer, the dendrimer is more hydrophobic, thus resultingin slower release profile of C₆ from the C₆-loaded dendrimer. Theconcentration of the dendrimer was 122 ug/ml.

FIGS. 14A-C. Uptake of dendrimer at concentration of 100 ug/ml by MDAcells at a temperature below the LCST (25° C.) and above the LCST (37°C.) of the dendrimers for 1 hour. A & B) The dendrimer was labeled withgreen FITC and the MDA cell nuclei were stained with blue DAPI. Confocalmicroscopy demonstrates that the dendrimers preferentially accumulateinto MDA cells at temperatures above the LCST (37° C.). Upper left, blueDAPI-stained nuclei; upper right, green FITC-dendrimer; lower left,phase/contrast; lower right, overlay. C) Flow cytometry analysisdemonstrates that significantly more dendrimer is internalized to MDAcells at a temperature above the LCST (37° C.) than below the LCST (25°C.) of the dendrimers. *p≦0.005.

FIG. 15A-B. C₆-Ceramide loaded dendrimers display anti-cancer effects invitro. A) In the presence of 5% FBS, C₆-loaded dendrimers permit thedelivery of ceramide to MDA cells resulting in C₆-induced cytotoxicitysimilar, if not better, to free C₆ administration in DMSO. B) C₆-loadeddendrimers result in significantly greater C₆-induced apoptosis thanfree administration of C₆ in DMSO. *p≦0.05.

FIG. 16. Structure of NIPAAM-co-PLLA-co-dextran hydrogels, wherein R isa —CONHCH₂CH═CH₂ or H, and m and n integers from about 1 to severalthousand. The NIPAAM segment can also have units of from about 10 toseveral thousand.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention provides a method and system for maximizing andtargeting the systemic delivery of growth-arresting pro-apoptotic,lipid-derived bioactive therapeutic compounds and/or hydrophobicchemotherapeutic agents and/or gene therapy agents to living cells of ananimal or human in need of such therapy, utilizing nanoscale assemblysystems, such as liposomes, resorbable and non-aggregatingnanoparticles, metal or semiconductor nanoparticles or polymericmaterials such as dendrimers or hydrogels, each of which exhibitimproved lipid solubility, cell permeability, an increased circulationhalf life and pharmacokinetic profile having improved tumor or vasculartargeting.

As used herein, “growth arresting” refers to living cells that are nolonger responsive to growth factors or cytokines released fromneighboring tissues. Moreover, “growth arrested” implies that the cellsdo not replicate their DNA and proliferate.

As used herein, pro-apoptotic refers to living cells or tumor tissuesthat undergo the process of programmed cell death.

As used herein, the phrase “lipid-derived” refers to substances that arethe metabolites of natural lipids found in biological membranes.

As used herein, the term “bioactive” refers to agents that transduceinformation and initiate a signaling cascade from the plasma membrane ofa cell to the nucleus where particular genes are either activated orinactivated resulting in a change to the phenotype of the cell (i.e.,growth arrest and/or apoptosis).

As used herein, the terms “nanoscale” and “nanosize” refer to a specialstate of subdivision implying that a particle has an average dimensionsmaller than approximately 300 nm and exhibits properties not normallyassociated with the bulk phase, e.g., quantum optical effects.

As used herein, the phrase, “hydrophobic chemotherapeutic agents” refersto small molecules, peptides, proteins, peptidomimetics, andlipidomimetics that are used as drugs to diminish cell proliferationand/or to induce cell apoptosis and are relatively insoluble in aqueousenvironments.

As used herein, the terms “nanocomposite particles” and “nanoparticles”are interchangeable.

As used herein, the term “agglomeration” refers to the formation of anaggregate (a cohesive mass consisting of particulate subunits) in asuspension through physical (van der Waals or, hydrophobic) orelectrostatic forces. The resulting structure is called an“agglomerate.”

As used herein, non-aggregating is the state of “dispersed”bioparticulates.

In particular, the present invention provides a system and method forsystemic, chronic or targeted delivery of a chemotherapeutic hydrophobiccompound to an animal or a human in need of such therapy that includes ananoscale assembly system and a growth-arresting, pro-apoptotic,lipid-derived bioactive compound or gene therapy agent. Nanoscaleassembly systems can include, without limitation, liposomes, resorbablenanoparticles that can be encapsulated by a calcium phospho-silicate(CPS) shell, or polymeric materials, such as dendrimers or hydrogels,that can be formed to be both bioresponsive (bio-smart) andbiodegradable.

The growth-arresting pro-apoptotic, lipid-derived bioactive compounds orgene therapy agents are delivered systemically via intravenous, catheterdelivery, infusion pumps, micro-spheres, or salves for treatingpathologies involving dysfunctional growth, such as cancer, neoplasm,arterial inflammatory disease, atherosclerosis, restenosis, vulnerableplaque or diabetes.

Examples of growth arresting pro-apoptotic, lipid-derived bioactivecompounds include, without limitation, physiological ceramides and/orderivatives, cell-permeable ceramides and/or derivatives that haveshort-chain fatty acids at the SN-2 position, consisting of 2-10 carbonunits, dimethyl sphingosine, trimethyl sphingosine, ether-linkeddiglycerides, ether-linked phosphatidic acids, sphingosines orsphinganines. Examples of gene therapy agents include, withoutlimitation, oligonucleotides, ribozymes, DNA-zymes, plasmids, antisenseor conventional Si-RNA or viral (AAV, AV or lenti) expressed SiRNA.

In one embodiment of the present invention, PEG-C₈ (750-5000 g/mol MW)and/or PEG-DSPE (2000-5000 g/mol MW) liposomes suitable for delivery ofhydrophobic bioactive lipids, proteins and therapeutic agents, referredto herein as “pegylated” liposomals, are formulated having one or moremembranes that are comprised of a growth-arresting lipid-derivedbioactive compound or a gene therapy agent and/or cholesterol. Liposomesthat are “pegylated,” with PEG-C₈ (750-5000 g/mol MW) and PEG-DSPE(2000-5000 g/mol MW) also known as “stealth” liposomes, can beformulated that are capable of evading clearance from the circulation bythe reticuloendothelial system (RES), and that can have binders attachedthereto, such as antibodies or receptor ligands to target specific cellsor tissues of the body. Liposomes, like other colloidal particles, areusually rapidly cleared from the circulation by the RES, primarily byKupfer cells of the liver and fixed macrophages of the spleen. The rateof liposome uptake by the RES is believed to be related to the processof opsonization or dysopsonization of the liposomes. Liposomaltherapeutic efficacy depends, therefore, on the ability to escaperecognition by the RES and thus remain in the circulation for prolongedperiods of time. The term “stealth” liposome, therefore, refers to thisevasive property and is conferred on liposomes whose membranes containbilayer-compatible species such as polyethylene glycol (PEG)-linkedlipids. “Stealth” or pegylated liposomes thus have the potential toimprove the hydrophilicity and bioavailability of drug releasingliposomes by evading the RES, and methods for liposome pegylationpreparation have been known for many years, as reported by Blume, G. etal. (Biochem. Biophys. Acta, 1029:91-97, 1990). Moreover, targeting canbe further achieved via the conjugation of particular targetingmoieties, such as antibodies and/or receptor ligands, to PEG, which willpromote the targeted accumulation into specific cells or tissues of thebody. Alternatively, the embodiment may contain cationic lipids, such asdioleoyl-1,2-diacyl-3-trimethylammonium-propane, used for effectivelydelivering negatively charged oligonucleotides. In addition, “fusogenic”liposomes can be formulated where the entire membrane of the liposomefuses with the cell membrane of the target site to deliver theconstituents and contents of the liposome therein. A fusogenic lipid isa destabilizing lipid that forms a hexagonal conformation in aqueoussolution, thus generating inverse micelles that bind to cell membranesvia an endocytotic or “fusogenic” process.

The liposomal vehicles of the present invention, therefore, amelioratesthe primary problems associated with systemic delivery of lipid-derivedbioactive compounds, such as C₆-ceramide, by preventing the bioactivelipid from precipitating out of solution so that it can be delivered tocells more effectively. Moreover, the present embodiment uses PEG-C8 tostabilize the lipid bilayer, allowing the liposome to containconcentrations of free bioactive C₆ ceramide of about at least 40 molarpercent. In addition, the embodiment utilizes the PEG-C8 as an integralcomponent of the liposome that contains the bioactive ceramide and/or ahydrophobic chemotherapeutic agent and/or a gene therapy agent.Moreover, PEG-C8 formulated liposomes ensures optimal intercalation andlocalization of the free ceramide into caveolin-rich lipid rafts, aprerequisite for membrane internalization and transfer to subcellularorganelles, such as the mitochondria, for subsequent induction ofapoptosis or programmed cell death of the targeted tissue or tumor.

Furthermore, the liposomes of the present invention can be applicablefor both local and systemic delivery of therapeutic ceramide analogues.For instance, it has been demonstrated that the local and directdelivery of C₆-ceramide from ceramide-coated balloons of embolectomycatheters limits neointimal hyperplasia (restenosis) in rabbits afterstretch injury. (Charles et al. Circ. Res. 2000 Aug. 18:87(4):282-8).Other groups have demonstrated that cell-permeable ceramide analogues inDMSO vehicle, delivered both intracistemally and intravenously, inducesa neuroprotective effect in rats following focal cerebral ischemia. Theclinical potential for the packaged delivery of C₆-ceramide withadditional therapeutic agents in liposomal vesicles is significant.Studies have shown that ceramide may act synergistically withchemotherapeutic agents, such as paclitaxel and fenretinide. Thus,combined delivery of chemotherapeutic agents in C₆-formulated liposomesmay further enhance apoptotic actions and at the same time diminish sideeffects by effectively lowering the concentration of each agent utilizedin the liposomal formulation. Moreover, targeted immunoliposomesconjugated with tumor-specific antibodies or receptor ligands may alsobenefit from C₆-ceramide incorporation.

The mechanism of ceramide involvement in the apoptotic program islargely unknown, although ceramide accumulation appears to be associatedwith a number of apoptotic hallmarks, such as poly(A)DP-ribosepolymerase (PARP) cleavage, DNA fragmentation, phosphatidylserineexposure and trypan blue uptake (Kolesnick, R. N. et al., Annu. Rev.Physiol., 60:643-665, 1998). Moreover, endogenous ceramide accumulateswithin mitochondrial membranes and, in part, induces cytochrome Crelease and resultant mitochondrial dysfunction and ultimatelyapoptosis. Ceramide can be generated through different metabolic routesin the cell. For example, it has been shown that there is astress-induced metabolic conversion of sphingomyelin into ceramide bythe enzyme sphingomyelinase in response to various treatments of cells,such as tumor necrosis factor-α(TNF-α), anti-Fas, serum withdrawal, andother agents. Additionally, ceramide can be generated through a de novosynthesis pathway, in which activation of serine palmitoyl transferaseand/or ceramide synthase may play a pivotal role (Garzotto, M. et al.,Cancer Res., 58:2260-2264, 1998). Exogenously added ceramides aregenerally able to mimic stress-induced apoptosis in a stereospecificmanner, and inhibition of the formation of ceramide has been shown insome cases to inhibit progression of apoptosis (Wiesner, D. A. et al.,J. Biol. Chem., 272:9868-9876, 1997). Exogenous ceramide intercalationand accumulation within caveolin-rich plasma membrane lipid rafts mayfacilitate internalization of these domains into subcellular organelles,including the mitochondria.

Apoptosis involves the orchestrated death of a cell and has been shownto be an important means by which organisms maintain homeostasis inproliferating tissues and systems, such as the immune system or ininflamed dysregulated cells or tissues as often observed in cancer,restenosis or atherosclerosis. (Frasca, L., et al., Crit. Rev. Immunol.,18:569-594, 1998). In fact, the loss of apoptosis control is a hallmarkof carcinogenesis. Ceramide analogues have been shown to induce cellularapoptosis in tumorogenic cells in vitro. However, to date, there are nostudies demonstrating the apoptotic and chemotherapeutic actions ofceramide in vivo, due to limited solubility upon systemic delivery. Theterm apoptosis often is used interchangeably with programmed cell death.It is distinguished from death by necrosis by the absence of anassociated inflammatory response. Apoptosis is characterized by theoccurrence of one or more cellular events that include loss ofmitochondrial integrity, nuclear condensation, membrane blebbing,chromatin fragmentation or loss of membrane integrity resulting inphosphatidylserine exposure and trypan blue uptake (Wyllie, A. H., J.Cell Biol., 73:189-197, 1997). Biochemical mechanisms by which each ofthese cellular characteristics are regulated remain largely unknown.However, it is believed that the activation of a family of cysteineproteases known as caspases plays an important role in the progressionof the apoptotic process (Thomberry, N. A. et al., Science,281:1312-1316, 1998). Within this caspase family, initiator caspases areactivated through an apoptotic stimulus and subsequently activatedownstream effector caspases. These effector caspases in turn have amultitude of intracellular substrates, among which are components thatare critically needed for cellular homeostasis. Cleavage of one or moreof these substrates disregulates cell function and promotes specificmorphological characteristics of the apoptotic program (Thomberry, N. A.et al., Science, 281:1312-1316, 1998). The mechanism by which PEG-C8stabilizes liposomes to allow up to about 40 per cent molar ratios offree bioactive ceramide available for membrane intercalation andinternalization and subsequent induction of apoptosis is a majorembodiment of the invention. Modulation of apoptotic processes bycompounds such as C₆-ceramide thus may offer valuable methods oftreatment.

In a further embodiment of the present invention, growth-arrestinglipid-derived bioactive compounds and/or gene therapy agents are loadedinto resorbable nanoparticles for drug and gene therapy having a calciumphospho-silicate (CPS) shell and a drug core. The resorbablenanoparticles of the present invention can deliver the hydrophobic lipidor protein drugs or gene therapeutic agents systemically to livingcells, which normally are not transportable through the circulation. Theresorbable nanoparticles can have a diameter ranging from 1 to 300 nm,preferably less than 50 nm and most preferably 20 nm or less.Nanoparticles having a diameter of about 20 nm or less are able to crossthe blood brain barrier (BBB), thus enabling the delivery of drugsdirectly into the central nervous system; a major advantage fortreatment of carcinogenic brain or neural lesions. These nanosystems arealso suitable for solid tumors, not limited to adenocarcinomas,melanomas, prostate, colon, lung (aerosol delivery) and breast tumors,as well as non-solid tumors such as leukemias The drug core either canbe delivered as a solid or in an aqueous solution. A schematic for thepreparation of the core-shell particles is shown in FIG. 1.

In particular, the method of synthesis for the resorbable nanoparticlesincludes a nonionic surfactant, such as poly(oxyethylene) nonylphenylether (IGEPAL™ 520 CO), or any other amphiphilic compound containing apolar head group and a non-polar tail, which is combined with water anda hydrophobic nonaqueous solvent, such as cyclohexane or iso-octanol inorder to form a reverse micelle structure. Growth-arresting,pro-apoptotic, lipid-derived compounds or gene therapy agents can besuspended in the aqueous phase as a solution, suspension or micellularmixture of water and drug or water and gene therapeutic agent. Theresulting reverse micelle containing the active agent within its core iscoated with an inorganic resorbable coating which biodegrades in aphysiological, i.e., isotonic, environment. CPS is an example of aresorbable coating, having the following composition:Ca_(x)(PO₄)_(y)zSiO₂, where 0.1≦x≦10 0.1, 0.1≦y≦10, and 0≦z≦10. Thecomposition can be adjusted to provide different rates of resorption ofthe shell in physiological environments, with higher silica and lowercalcium and phosphate concentrations resulting in a slower resorptionrate.

A key feature of the synthesis of the resorbable nanoparticles is theproper dispersion of the nanoparticles in a liquid medium. Suitableliquids can include, without limitation, deionized water, salinesolution, water-ethanol mixtures or other liquid-suspending mediasuitable for a physiological environment and/or any additionalprocessing steps, e.g., granulation processes such as spray-drying priorto tablet formulation for oral delivery. This is accomplished first bywashing the nanoparticles in order to remove excess amphiphilic compoundand any other ions or additives to ensure that optimal dispersion of thenanoparticles is achieved. Additionally, it is necessary to concentratethe suspension during washing to produce a suspension at a high enoughconcentration in order to deliver the drug or gene therapy agent at asufficient dosage. This is achieved using size exclusion highperformance liquid chromatography (SEC) modified specifically for thesilicate-containing shell nanoparticles. Such modifications arenecessary to prevent solid bridge formation between the contactingnanoparticles that results in persistant agglomeration. The size of theprimary nanoparticles produced according to the method of the presentinvention can range between about 1.0 to 300 nm. This procedure preventsagglomeration of the nanoparticles, which can be well over 1 micron, asmeasured by particle size distribution measurement techniques, such asquasi-elastic light scattering or centrifugation or sedimentation withoptical density determination. Thus, nanoparticle suspensions that arenot processed as described herein can result in significantly alteredflow units due to the agglomeration during the washing and recoverysteps than what would be the case for primary size nanoparticles.

The critical modifications of the SEC collection and washing steps ofthe nanoparticles includes using shorter elution columns that containmicroporous silica particles having a diameter of about 20 microns, aswell as chemical modifications. Chemical modifications include, withoutlimitation, adding ethanol or any other suitable alcohol to the reversemicelle suspension after synthesis of the nanoparticles to produce ahomogeneous nanoparticle suspension. The substitution of alcohol inplace of the typical use of water or acetone prevents the formation ofagglomerated masses of nanoparticles having flow units much greater thanprimary nanoparticle size flow units. Another critically importantchemical modification is the attachment of an organic or inorganicdispersant which acts on the surface of the nanoparticles to provide anelectrosteric layer that prevents the nanoparticles from persistentagglomeration. Suitable organic dispersants include, without limitation,citric acid, tartaric acid or acetic acid. Suitable metal-organicdispersants include, without limitation, alkylamine silane couplingagents, such as aminopropyltrichlorosilane;3-aminopropyltrimethoxysilane (APS); 3-aminopropylsilsesquioxane);3-glycidoxypropyl-trimethoxysilane (GPS);trimethoxysilyl-propyldiethylenetriamine (DETA);3-trimethoxysilypropylsuccinic anhydride; and alkylcarboxylic acidsilane coupling agents, such as amide-linked carboxyl groups. Suitableionic dispersants include, without limitation, excess calcium, phosphateor pyrophosphate.

Furthermore, it is necessary that the microporous silica particles thatare used to pack the SEC column be surface-treated with the identicaldispersing agent in order to produce an electrosteric barrier thatprevents the nanoparticles from adhering to the microporous silicasurfaces during passage through the column. It also is necessary tocontrol pH levels during the SEC nanoparticle concentration and washingsteps. Thus, acids, such as, without limitation, nitric acid, aceticacid or hydrochloric acid; or bases such as, without limitation, sodiumhydroxide or potassium hydroxide, are added as needed in order tomaintain pH levels within a range of between about 6 and 8. With a pHgreater than 8, the charge on the surface of the nanoparticles is toolow and agglomeration results. With a pH less than 6, the concentrationof acid or base is too high and the resultant ionic strength can causeagglomeration during the washing step. Ceramides and other lipid-derivedbioactive mediators are resistant to all of these acidic or alkalineprocedures.

Additionally, a carbodiimide-mediated polyethylene glycol (PEG) couplingagent can be attached to the alkylamine silane or alkylcarboxylic acidcoupling agent to further ensure the dispersed state of thenanoparticles in vivo and to provide an attachment point for binders,such as antibodies or ligands for expressed receptors, onto the PEGcoupling agent, thus enabling intracellular drug delivery of theceramide enriched or encapsulated nanoparticles to targeting tumorspecific sites.

In another embodiment of the present invention, the nanoscale assemblysystem is comprised of polymeric material, such as dendrimers orhydrogels, which are loaded with growth-arresting, pro-apoptotic,lipid-derived bioactive compounds and/or gene therapy agents. Thedendrimers or hydrogels are individual polymers that are combined toform materials that are both bio-smart, i.e. respond to stimuli, andbiodegradable in vivo. It was discovered, after investigation andexperimentation, that materials which combine a smart segment with adegradable hydrophobic and/or hydrophilic segment can be used for drugdelivery. A segment is considered to be a covalently bound portion ofthe material and can have a plurality of polymerized units. For example,a segment can include several polymerized monomer units up to aboutseveral thousand polymerized monomer units. These segments can have anylength and any molecular weight, however it is preferred that eachsegment has a molecular weight that is roughly large enough toapproximate a desired property expected for that polymer segment.

The polymeric material when used alone is limited by sub-optimal ornon-biodegradability. Thus, combining a smart polymer segment with abiodegradable polymer segment results in a material considerably moreversatile than the individual materials. By combining both bioresponsiveand biodegradable polymers, a drug delivery system is fashioned which isboth biodegradable and responsive to physiological stimuli. Inparticular, a multifunctional polymeric material is provided comprisinga smart segment and a biodegradable segment, wherein the biodegradablesegment includes a hydrophobic segment (suitable for binding orinteracting with a chemotherapeutic agent) and a hydrophilic segment.

A number of natural and synthetic biodegradable polymers are known. Somehave been studied, including polyesters, such as polylactides (PLA),poly(L-lactic acid), poly(D,L-lactic acid), poly(lactide-co-glycolides)(PLGA), biotinylated poly(ethylene glycol-block-lactic acid),poly(alkylcyanoacrylates) and poly(epsilon-caprolactone);polyanhydrides, such as poly(bis(p-carboxyphenoxy) propane-sebacic acid)(PCPP-SA), polyorthoesters, polyphosphoesters, polyphosphazenes,polyurethanes, and poly(amino acids), polysaccharides, such as dextran,in the forms of microcapsules, microparticles, nanoparticles, hydrogelsand micelles. All such biodegradable polymers are contemplated in thepresent invention as segments of a multifunctional material.

The forgoing polymers degrade by hydrolytic or enzymatic cleavage of thebackbone and, hence, are non-toxic and non-inflammatory after drugdepletion. The degradation properties of the polymers depend on theirchemical composition, tacticity, crystallinity, molar mass, morphology,size and shape, and also pH and temperature. The chemical and physicalproperties of biodegradable polymers are known to influence the drugrelease patterns, and the release kinetics of the loaded drugs arecontrolled by both drug diffusion and polymer degradation.

Another approach to affect the degradation rates of dendrimers orhydrogels include coating or grafting hydrophobic materials, e.g. PLAand PLGA micro/nanoparticles, with poly-L-lysine (PLL) due to the PLL'scharge, hydrophilicity and targeting capability. For example, it hasbeen shown that microparticles composed of poly(L-lacticacid-co-L-lysine) grafted with PLL have significantly increased releaserates of rhodamine B compared to those without the PLL side chains.Furthermore, PLGA grafted with PLL micelles display 10 times highertransfection efficiency and 5 times less cytotoxicity than PLL. Thepresent invention contemplates the use of such coating and graftingtechniques in providing hydrophilic-hydrophobic degradable segments.

Dendrimers are defined by regular, highly branched segments leading to arelatively monodisperse, tree-like or generational structure. Dendrimerspossess three distinguishing architectural features: the core; theinterior area containing branch upon branch of repeat units orgenerations with radial connectivity to the core; and an exterior orsurface region of terminal moieties attached to the outermostgeneration. A dendrimer can be defined into a multitude of structures bytuning these three architectural components. Dendrimers that are highlybranched and reactive three-dimensional macromolecules have becomeincreasingly important in biomedical applications due to their highdegree of molecular uniformity, narrow molecular weight distribution,specific size and intriguing structural properties such as internalvoids and cavities, and a highly functional terminal surface. Thespatially arranged functional groups can react with a variety ofmolecules, for example, hydrophilic molecules such as PEO (polyethyleneoxide or PEG) to increase their blood circulation times, contrast agentsfor use in magnetic resonance imaging (MRI), and targeting molecules tolocalize to desired tissue sites.

Currently available dendrimers contain benzyl ether, propyleneimine,amidoamine, L-lysine, ester and carbosilane dendritic segments. Amongthem, cationic polyamidoamine (PAMAM) dendrimers have been widelystudied and were reported to mediate high levels of gene transfection ina wide variety of cells, depending on the dendrimer-DNA ratio, the sizeand especially the flexibility of the dendrimers. PAMAM dendrimers areconsidered targeted delivery systems, and can enhance accumulationwithin certain tumor microvasculature, increase extravasation into tumortissue. Poly(L-lysine) (PLL) dendrimer is another polycationic dendrimercontaining a large number of surface amines and considered to be capableof the electrostatic interaction with polyanions, such as nucleic acids,proteoglycans found in extracellular matrix and phospholipids of thecell membrane. These polymers can localize drugs, includinglipid-derived bioactive growth arresting, pro-apoptotic metabolites oragents to the targeted membranes.

However, polycationic dendrimers still have in vivo toxicity problemsand are resistant to degradation in the body and are thus less suitablefor drug delivery. To improve the cytotoxicity of PAMAM dendrimers, thecationic amine terminal groups of the dendrimers can be replaced withanionic carboxylate terminal groups. The present inventive materialsaddress some of the disadvantages of dendrimer structures prepared fromindividual components by combining smart and degradable segments asarms, branches, or dendrons of a dendrimeric structure. Such dendrimericmaterials can be prepared by coupling a thermoresponsive polymer segmentwith a biodegradable polymer segment in a chemical bond formingreaction.

Dendrimers also can be prepared as nano-sized particles. It is believedthat particles having a size of about 1 nm to 1000 nm hold a significantadvantage in transporting and targeting drugs to inflamed, proliferativeor transformed tissues. Drugs are loaded into the nano-sized dendrimersby adsorption, entrapment and covalent attachment, and released from thenano-sized dendrimers by desorption, diffusion, polymer erosion or somecombination of any or all the above mechanisms. In vitro and in vivoexperiments show that nano-sized dendrimers can have long bloodcirculation times and a low RES uptake when they are stabilized bydextran and coated with polysorbate 80. The nano-sized dendrimers may beable to interact with the blood vessel or solid tumor cells, and then betaken up by these cells by endocytosis. Dendrimers are believed,therefore, to have a great potency to deliver drugs to tumorigenic orinflamed/proliferative tissues due to increased circulatory half-life.

Recent advances in nanotechnology offer enormous potential forcontrolled delivery and targeted release of hydrophobic therapeutics.The nanoscale dendrimeric assembly system of the present invention isboth responsive to temperature stimuli and hydrolytically biodegradable,allowing for the targeted and sustained delivery of C₆ to solid tumors.It has been demonstrated that C₆ can be loaded intotemperature-sensitive, “smart” dendrimers, and that this drug-polymercomplex can effectively inhibit the proliferation, as well as induceapoptosis, of MDA estrogen-negative breast cancer cells. The applicationof acute local hyperthermia, via a heat pack, infrared or ultrasound, tothe area of a solid tumor will trigger the release of C₆ into diseasedtissue. Thus, thermo-responsive nanoscale dendrimers can serve as anoptimal solution for targeted and controlled delivery of therapeuticagents, including ceramide, to solid tumor tissue, a concept coined as“physiological hyperthermic drug delivery.”

Dendrimers in Drug Delivery.

Liposomal drug delivery technology is slowly being eclipsed by moreadvanced drug delivery systems that incorporate polymer chemistrytechnology in order to engineer stable nanoparticles with a dynamicarray of drug delivery advantages. For instance, polymeric nanoparticlesare capable of prolonged bioavailability, diseased cell targeting, andbioresponsive and controlled drug release (17). Drugs are loaded intothe polymeric nanoparticles by adsorption, entrapment and covalentattachment, and released from the nanoparticles by desorption,diffusion, polymer erosion or some combination of any or all mechanisms.Dendrimers, highly branched and reactive three-dimensionalnanoparticles, are suitable for biomedical applications due to theirhigh degree of molecular uniformity, narrow molecular weightdistribution, specific size and intriguing structural properties such asinternal voids and cavities, and a highly functional terminal surface.The dendrimers of the present invention are comprised of a polycationicpolymer (poly(L-lysine), PLL), a biodegradable polymer (poly(L-lacticacid), PLLA), and a thermo-responsive polymer(poly(N-isopropylacrylamide), PNIPAAM). Hydrophobic agents, such as C₆,are loaded into the dendrimer in concentrations up to 1000 mg/ml, byhydrophobic-hydrophobic interactions with the PLLA. Incorporation ofresponsive polymers with biodegradable polymers has advantages inachieving sustained release of drugs in response to a physiologicalstimuli, such as temperature.

Smart or responsive polymers are responsive to physical, chemical, orbiological stimuli, such as temperature, solvent composition, pH, ionicstrength, pressure, electric field, light and metabolites. Amongthermo-responsive polymers, poly(N-isopropylacrylamide) (PNIPAAM) hasbeen extensively used for controlled drug delivery, since it exhibits aunique solubility transition at the lower critical solution temperature(LCST) in an aqueous solution in the vicinity of 32° C. It expands andswells when cooled below the LCST, and it shrinks and collapses whenheated above the LCST. The LCST of PNIPAAM can be manipulated forcontrolling the loading and the release of drugs by incorporatinghydrophobic and hydrophilic units, and crosslinkers into PNIPAAM.

Importantly, PNIPAAM-based polymers can be used as reversible targetingmoieties for site-specific drug delivery. In the present invention, thepolymers are designed with the LCST between 37° C. and 42° C. At a bodytemperature of 37° C., below the LCST, the polymers are soluble in thephysiological fluids, evade the body's reticulo-endothelial system (RES)and increase the loaded drugs' blood circulation time. When thetemperature is increased to 42° C. via, without limitation, localultrasound, infrared and/or heat patches, which is a temperature higherthan the LCST of polymers at the targeting site, the polymers accumulateat the targeting site and release therapeutic drugs with high localconcentrations. It has been shown that systemic injection ofpoly(NIPAAM-co-acrylamide) with a LCST of 40° C. in mice accumulated thecopolymer at solid tumors by local hyperthermia at a 2 fold greaterdegree than that for heated and unheated control groups. Additionally,biodegradable and thermoresponsive micelles composed ofpoly(N-isopropylacrylamide-co-N,N-dimethylacrylamide)-b-poly(D,L-lactide)have been prepared with a LCST of 40° C. It has been shown that thecytotoxicity of the anticancer drug adriamycin loaded in the micellesagainst bovine aorta endothelial cells was higher than that of freeadriamycin above the LCST at 42.5° C. due to accelerated uptake of themicelles by the cells. The dendritic nanoparticles of the presentinvention expand upon the thermo-responsive properties of PNIPAAM, asthe polymer is covalently linked to polycationic PLL (hydrophilicity andstability) and biodegradable PLLA (hydrophobic, controlled-release).Moreover, this multi-functional dendrimer can be loaded with thepro-apoptotic lipid C₆ to induce breast cancer cell apoptosis.

Hydrogels are three-dimensional crosslinked polymer networks that swellin an aqueous environment by absorbing large amounts of water whilemaintaining their structure. Due to their high water content,biocompatibility, and unique mechanical properties, hydrogels haveattracted wide interests in biomedical applications such as drugdelivery and tissue engineering. The environmentally-sensitive hydrogelsof the present invention can control drug release by changing theirstructures in response to environmental stimuli, such as temperature,pH, electrical signal, ionic strength, etc. Covalently andnon-covalently (physically) crosslinked temperature-sensitive,biodegradable gels are preferred materials as hydrogels.

In particular, hydrogels are prepared as copolymeric networks composedof N-isopropylacrylamide (NIPAAM) or a derivative thereof as a smart orresponsive component; poly(L-lactic acid) (PLLA) or a derivative thereofas a hydrolytically degradable and hydrophobic component; and dextran ora derivative thereof as an enzymatically degradable and hydrophiliccomponent. The components or segments can be of any length includingfrom about 3 monomer units to about 10,000 monomer units, e.g. about 3to 5,000 units. The material or segments can further comprise othermonomer units to adjust the materials properties. For example, thehydrogel can also include anionic (acrylic acid) and cationic (acrylicamine) units for increasing pH and ionic strength sensitivity of thegel.

PNIPAAM-PLLA-dextran hydrogels are thermo-responsive showing a lowercritical solution temperature (LCST) at approximately 32° C., and theirswelling properties strongly depend on temperature changes, the balanceof the hydrophilic/hydrophobic components and the degradation of thePLLA component. The degradation of the hydrogels caused by hydrolyticcleavage of ester bonds in PLLA component, is faster at 25° C., belowthe LCST than at 37° C., above the LCST, as determined by ATR-FTIR andweight loss measurement.

It is stated, without being bound by the theory, that when therapeuticcompounds, such as growth-arresting, pro-apoptotic, lipid-derivedbioactive compounds and/or gene therapy agents, are incorporated intonanoscale assembly systems, such as liposomes, resorbable,non-aggregating nanoparticles, or polymeric materials, such asdendrimers or hydrogels, their systemic delivery to cancerous cells isaugmented, thus enhancing greatly the potency of the growth-arresting,pre-apoptotic compounds. Additionally, ceramide-enriched or encapsulatednanotechnology can be engineered to deliver lower doses of otherhydrophobic chemotherapeutic agents or gene therapies in a combinationtherapy to achieve higher efficacy with diminished side effects.

The following examples are intended to further illustrate certainpreferred embodiments of the invention and are not limiting in nature.Those skilled in the art will recognize, or be able to ascertain, usingno more than routine experimentation, numerous equivalents to thespecific substances and procedures described herein.

EXAMPLE 1 C₆-Ceramide-induced Apoptosis of Breast Cancer Cells viaLiposomal Delivery

Materials and Cell Culture

Egg phosphatidylcholine (EPC), dioleoyl phosphatidylethanolamine (DOPE),dioleoyl phosphatidylcholine (DOPC), cholesterol (CH),polyethyleneglycol (2000-5000)-distearoyl phosphatidylethanolamine(PEG-DSPE), D-erythro-hexanoyl-sphingosine (C₆-ceramide),polyethyleneglycol-750-C₈-ceramide (PEG-C8),dioleoyl-1,2-diacyl-3-trimethyl-ammonium-propane (DOTAP) were purchasedfrom Avanti Polar Lipids (Alabaster, Ala.). Di-hydro-erythrohexanoyl-sphingosine (DHC₆) was purchased from Biomol (Plymouth Meeting,Pa.). [³H]-C₆ was obtained from ARC (St. Louis, Mo.), [³H]-thymidine waspurchased from ICN (Costa Mesa, Calif.) and cholesteryl-1,2-³H(N)hexadecyl ether ([³H]-CHE) was obtained from PerkinElmer (Boston,Mass.). Silica gel 60 thin layer chromatography plates were purchasedfrom EMD Chemicals (Gibbstown, N.J.). Formvar/carbon-coated 400 meshcopper grids were purchased from Electron Microscopy Sciences (FortWashington, Pa.), and poly-L-lysine was obtained from Sigma (St Louis,Mo.).

Antibodies specific for phosphorylated-Akt (pAkt) and Akt-1,2,3 werepurchased from Cell Signaling (Beverly, Mass.). Insulin-like GrowthFactor-1 (IGF-1) was obtained from CalBiochem (San Diego, Calif.). ForWestern blotting, 4%-12% pre-casted SDS-PAGE gradient gels were obtainedfrom Invitrogen (Carlsbad, Calif.) and ECL reagent from Amersham(Piscataway, N.J.). The TUNEL Apoptosis Detection Kit was obtained fromUpState Biotechnology (Waltham, Mass.). The Vybrant Apoptosis Assay Kit#3 was purchased from Molecular Probes (Eugene, Oreg.), and the Apo-ONEHomogeneous Caspase-3/7 Assay was obtained from Promega (Madison, Wis.).RNAse was purchased from Roche (Indianapolis, Ind.) and propidium iodide(PI) from Sigma (St. Louis, Mo.). Human MDA-MB-231 (MDA) breastadenocarcinoma cells were obtained from ATCC (Manassas, Va.) and grownat 37° C. in RPMI 1640 supplemented with 10% FBS. This cell line is ahighly aggressive metastatic, estrogen receptor-negative, model of humanbreast cancer.

Liposome Formulation and Extrusion

Lipids were formulated and tested for their ability to incorporate C6into liposomal drug delivery vesicles. Briefly, lipids, dissolved inchloroform (CHCl3), were combined in specific molar ratios, dried undera stream of nitrogen above lipid transition temperatures, and hydratedwith sterile phosphate-buffered saline (PBS). The resulting solutionunderwent sonication for 2 min followed by extrusion through 100 nmpolycarbonate membranes. Incorporation efficiency was determined byincorporating trace amounts of [3H]C6 in the formulation, extractingconstituent lipids in CHCl3/MeOH (2:1), and comparing radioactivitybefore and after extrusion using a scintillation counter. Formulationsfor in vivo administration comprised ofDSPC:DOPE:DSPE-PEG(5000):C8-Ceramide-PEG(750):C6-Ceramide(3.75:1.75:0.75:0.75:3.0, molar ratios). The addition of PEG(750)-C₈allows for up to 40 molar percent C₆-ceramide. The bioactivity of thesepegylated formulations were confirmed in 410.4 mammary adenocarcinomacells. The composition of formulated liposomes was validated byextracting constituent lipids in chloroform/methanol (2:1), followed byresolution on preheated silica gel 60 thin layer chromatography (TLC)plates using a CHCl3/MeOH/ddH2O (60:25:4) solvent system. Lipids werevisualized in an iodine chamber. Transmission electron microscopy (TEM)was utilized to characterize the size and morphology of the formulatedliposomes.

Transmission Electron Microscopy (TEM)

In order to characterize the size and morphology of the formulatedliposomes, TEM was utilized. Initially, formvar carbon-coated 400 meshcopper grids were coated with poly-L-lysine for 10 minutes, in order topromote vesicular binding to the hydrophobic grids. Liposomal sampleswere next applied to the dried grids and allowed to adhere for 5minutes. Negative staining was performed by applying 1% phosphotungsticacid (pH 7.0) to the dried grid for an additional 5 minutes. The samplewas observed at 21,500× magnification with an accelerating voltage of 60kV.

TEM analysis confirmed that C6-incorporated liposomal vehicles wereproduced with a homogeneous size distribution between 85 and 140 nm indiameter for all formulations (FIG. 2A). FIG. 2B illustrates the averagesize of the liposomal formulations. With the incorporation of traceamounts of [3H]C6 into conventional formulations, we observed that therewas no significant loss of ceramide during the extrusion process (FIG.2C). Additionally, lipid extracts from conventional liposomes were runon TLC plates, confirming that there was no visual diminution of lipidconstituents during the extrusion process (FIG. 2D).

In Vitro Pharmacokinetics

Trace amounts of [³H]C₆ were incorporated into liposomal formulations toquantify the amount of liposomal delivery compared to nonliposomaladministration. Human MDA-MB-231 (MDA) breast adenocarcinoma cells wereseeded at 3.5×10⁴ cells/well in 24-well plates and grown overnight inmedia containing 10% FBS. Cells were then treated with liposomal ornonliposomal C₆ containing trace amounts of either [³H]C₆ or [³H]CHE inmedia supplemented with 1% FBS for various time intervals. Liposomal C₆was added directly to cell media, and nonliposomal C₆ was added indimethylsulfoxide (DMSO) vehicle to a final concentration of ≦0.1%(v/v). At the indicated time points, the media was removed, and cellswere washed once with cold PBS to dissociateliposome/membrane-nonspecific interactions. The cells were thensolubilized with 1% SDS, and either [³H]C₆ or [³H]CHE accumulation intoMDA cells was assessed with a scintillation counter.

Results showed that liposomal formulations delivered C₆ more effectivelyand efficiently than nonliposomal administration of C₆ in the presenceof 1% FBS (FIG. 3A). Cationic liposomal delivery resulted in a 2-foldincrease in ceramide accumulation by MDA cells, with a maximalaccumulation observed at approximately 16 h. Conventional and pegylatedliposomes were observed to have similar in vitro pharmacokineticprofiles.

The mechanism by which C₆ is released or transferred from liposomalvehicles into cellular membranes then was investigated. Using anontransferable cholesterol lipid marker, [³H]CHE, as a probe forliposome/cell membrane association, liposomes were tagged with either[³H]C₆ or [³H]CHE and incubated with MDA cells for the indicated timeperiods. As shown in FIG. 3B, liposomes mediated the transfer of C₆, butnot cholesterol, from drug vehicle to cellular membrane. Furthermore, asC₆ accumulation increased over time, CHE accumulation failed tosignificantly increase above background levels. The disparity betweenceramide and cholesterol accumulation also is observed in adose-dependent manner (FIG. 3C). This suggests that C₆ is delivered vialipid transfer processes that permit C₆ to partition out of theliposomal layer into the plasma membrane bilayer without associatedliposome/cell membrane fusion.

[³H]-Thymidine Cell Proliferation

To determine the utility of conventional lipid formulations for thedelivery of short-chain ceramide to MDA cells, a [³H]thymidineproliferation assay was performed. Briefly, MDA cells were seeded at3.5×10⁴ cells/well in 24-well plates and grown overnight prior to 24 hof serum starvation. At hour 12 of serum starvation, cells were treatedwith liposomal or nonliposomal C₆ for the remainder of serum starvation.Following serum starvation, media was then supplemented with FBS (10%final concentration) for an additional 12 h, and cellular proliferationwas assayed with the addition of 0.5 mCi/ml [³H]thymidine for the final4 h of treatment. Cells were washed once with cold PBS and then twicewith 10% trichloric acetic acid for 10 min. Cells were solubilized with0.3 N NaOH, and [³H]thymidine incorporation into acid-insoluble DNA wasassessed with a scintillation counter.

As shown in FIG. 4A, a conventional liposome containing eggphosphatidylcholine (EPC) and cholesterol (CH) (solid lines)supplemented with C₆ displayed a significant dose-dependent inhibitionof MDA cell proliferation. The addition of a vesicle-destabilizinglipid, dioleoyl phosphatidylethanolamine (DOPE), into a conventionalformulation also enhanced the bioactivity of C₆. MDA cells, in thepresence of 10% FBS for 12 h of treatment, were completelygrowth-inhibited when treated with liposomal C₆ at 25 μM or greater. Thedelivery of C₆ in liposomal formulations reduced the IC50 approximately3-fold, decreasing from 15 to 5 μM, nonliposomal to liposomal,respectively. These conventional formulations displayed an improveddose-response inhibition of growth in MDA cells compared withnonliposomal administration of C₆ in DMSO vehicle (dashed line, opencircle), indicating improved potency and efficacy. Liposomes without C₆(Ghost; dashed line, open square) as well as PBS controls did notdisplay significant growth inhibition, implicating C₆ as the onlybioactive agent. This study demonstrated that that C₆-formulatedconventional liposomes were more effective as an antiproliferative thanfreely administered C₆.

C₆-incorporation into cationic lipid formulations next were investigated(FIG. 4B). Even though Ghost cationic liposomes (dashed line, opentriangles) formulated with a positively charged lipid,dioleoyl-1,2-diacyl-3-trimethylammonium-propane (DOTAP), enhanced MDAcell proliferation alone, C₆-incorporated cationic liposomes (solidline, open triangle) dose-dependently reduced MDA cell proliferation.This cationic formulation was more effective than nonliposomal C₆administration (dashed line, open circle) but not as effective as aconventional formulation (solid line, open square). This indicates thatcationic liposome formulations could also be used to deliver bioactiveceramide to dose-dependently inhibit cell proliferation.

The role of pegylated lipid to further enhance the bioactivity of C₆next was investigated. C₈-ceramide (PEG-C8) was chosen because of itsadditional potential benefit to promote liposome/membrane fusion.Additionally, the inclusion of PEG-C8 is known to facilitatetime-release properties of liposomal bilayers, with the added benefit ofbioavailability extension. Moreover, the present embodiment uses PEG C8to stabilize the lipid bilayer, allowing the liposome to containconcentrations of free bioactive C₆ ceramide up to at least 30 molarpercent. In addition, the embodiment utilizes the PEG C8 as an integralcomponent of the liposome that contains the bioactive ceramide and/or ahydrophobic chemotherapeutic agent and/or a gene therapy agent.Moreover, PEG-C8 formulated liposomes ensures optimal intercalation andlocalization of the free ceramide into caveolin-rich lipid rafts, apre-requisite for membrane internalization and transfer to subcellularorganelles including the mitochondria for subsequent induction ofapoptosis or programmed cell death of the targeted tissue or tumor.

Pegylated liposomes did not markedly effect MDA cell proliferation,demonstrating that PEG-C8 is biochemically inert. However,C₆-incorporated pegylated liposomes were as if not more, effective atinhibiting proliferation as conventional liposomes at 10 and 25 μM (FIG.4C). This indicated that a pegylated liposomal formulation designed forsystemic drug delivery also is an effective vehicle for C₆-mediatedinhibition of MDA cell proliferation. Taken together, C₆ delivered inmultiple liposomal formulations displays an improved dose-responseinhibition of growth compared with nonliposomal C₆, indicating improvedpotency and efficacy.

MTS Cytotoxicity Assay

To assess the in vitro efficacy of pegylated formulations used for invivo studies, we tested the formulations on murine 410.4 mammaryadenocarcinoma cells. The 410.4 cells were plated in 96-well plates andtreated with pegylated liposomal or free C6 for 24 hours in culturemedia supplemented with FBS to 1%. Cytotoxicity was assessed using thePromega Cell Titer Proliferation Kit (Promega) according to themanufacturers instructions. Pegulated liposomal-C₆ delivery results inenhanced cellular toxicity. The administration of liposomal-C6formulations lowers the IC50 of C6, compared to free administration ofC6 in DMSO vehicle (FIG. 5). These data indicate a 35-40% reduction inthe IC₅₀ of C₆-ceramide when delivered in PEG-C₈ liposomal formulations.Treatments were performed in the presence of 1% FBS for 24 hours.

Caspase Assay

Apoptosis is associated with the up-regulation of caspase activity, thuscaspase-3/7 activity following treatment of MDA cells with pegylatedliposomes was assessed. Briefly, MDA cells were seeded to a density of6.0×10³ cells/well in 96-well plates and grown for 48 h in culture mediacontaining 10% FBS. Cells were then treated with liposomal ornonliposomal C₆ for 24 h in media containing 1% FBS. Caspase-3/7enzymatic activity levels were measured using the Apo-ONE homogeneouscaspase-3/7 assay (Promega, Madison, Wis.) according to standardprotocol known in the art

The results showed that MDA cells treated with pegylated liposomalceramide displayed significantly greater caspase-3/7 activity than cellstreated with nonliposomal ceramide (FIG. 6). No significant change incaspase-3/7 activity was observed with Ghost treatments. Taken together,these results indicate that C₆-formulated liposomes were more effectivethan nonliposomal administration of C₆, resulting in significantinhibition of MDA cell proliferation and eventual apoptotic death.

Apoptosis Detection

An investigation to determine whether C₆-dependent growth inhibitioncorrelates with enhanced apoptosis was undertaken. To confirm that C₆delivery leads to MDA cell apoptosis, TUNEL analysis (UpstateBiotechnology, Lake Placid, N.Y.) was performed, which stains cleavedDNA, a hallmark of cellular apoptosis (FIG. 7A).

TUNEL staining of cycling, serum-fed MDA cells treated with liposomaland nonliposomal C₆ demonstrated no DNA fragmentation at 8 h. Liposomaland nonliposomal C₆ treatment induced DNA fragmentation in a similarmanner to the Dnase-positive control. Staining of cleaved 3′-OH DNA wasobserved at 16 h of treatment, a time point consistent with the in vitropharmacokinetic profile of C₆ delivery. No apoptosis was observed withthe Ghost formulation.

To quantitate the C₆-induced apoptosis, annexin V staining of treatedcycling MDA cells and flow cytometry analysis of the annexin V-stainedcells, using the Vybrant apoptosis assay kit (Molecular Probes, Eugene,Oreg.), was performed.

Following a 24 h treatment, pegylated liposomal C₆ induced asignificantly greater amount of annexin V staining compared withnonliposomal C₆, whereas the Ghost formulation had no effect (FIG. 7B).

Assessment of Activated AKT, A Pro-Survival Kinase

Ceramide-regulated Akt signaling pathways in MDA cells treated withliposomal and C₆ nonliposomal formulations was investigated usingWestern blot analysis. Briefly, MDA cells were seeded at 4.0×10⁵cells/well in 60 mm plates and grown overnight, prior to 24-hour serumstarvation. At hour 16 of serum starvation, cells were treated withliposomal or non-liposomal C₆ for the remainder of serum starvation. Athour 24 of serum starvation, IGF-1 (20 ng/ml) was added to cell mediafor a 15 minute period. Cells were washed once with cold PBS followed bythe addition of 150 ul of cold lysis buffer (1% Triton X-100, 20 mMTris, 150 mM NaCl, 1 mM EDTA, 1 mM EGTA, 2.5 mM Na₄P₂O₇, 1 mMβ-glycerolphosphate, 1 mM Na₃VO₄, 1 μg/ml leupeptin in ddH₂O, pH 7.5) onice. Cells were lysed for 15 minutes on ice, cell lysate was harvested,and centrifuged at 15,000×G for 15 minutes. 35 μg of protein were loadedin 4%-12% pre-casted SDS-PAGE gradient gels and probed for pAkt. Blotwere stripped and re-probed for Akt-1,2,3 to demonstrate equal loading.Protein bands were visualized using ECL chemiluminescence and quantifiedby densitometry. The results showed that pegylated liposomal C₆ was moreeffective at reducing IGF-1-stimulated pAkt levels than was nonliposomalceramide, whereas the Ghost formulation had no effect on pAkt levels(FIGS. 8A and 8B). Liposomal dihydro-erythro-hexanoyl-sphingosine (DHC₆)also displayed inhibition of IGF-1-stimulated Akt phosphorylationcompared with nonliposomal DHC₆. Eight hours of C₆ treatment wasselected, as this time point corresponded to near maximal accumulationof C₆ into MDA cells. This study supports showed that liposomal C₆induced cell growth inhibition and apoptosis through long-terminhibition of Akt signaling cascades.

Confocal Studies

In order to verify cell accumulation of C₆ into 410.4 murine mammaryadenocarcinoma cells, we administered liposomal-C₆ formulations with 10molar % NBD-C₆ as a marker for C₆. Cells were counter-stained with DAPI(nuclei) and MitroTraker Red (mitochondria) for reference. C₆ deliverywas evaluated by confocal microscopy at a magnification of 60×. Confocalmicroscopic images of NBD-C₆ delivery to cells from liposomal vesicles(FIG. 9A). NBD-C₆ (Green) co-localized with mitochondria(MitoTraker-Red); blue stained represents DAPI-stained nuclei.

Sucrose Gradient

The incorporation of trace amounts of [³H]-C₆ into liposomalformulations was utilized to assess a time-dependant cellularaccumulation of C₆ in both total cells and caveolae-enriched lipidrafts. As these lipid rafts are believed to facilitate the signaltransduction of multiple pathways, including ceramide, theadministration of liposomal-C₆ should facilitate the accumulation of C₆into lipid rafts. In order to evaluate this phenomenon, a sucrosegradient (5%, 35%, and 45% sucrose) of cellular lysate was performed inorder to isolate the caveolae-enriched lipid rafts. Equal aliquots ofeach 1 ml fraction of the gradient (total of 12 fractions) were removedand counted using a scintillation counter. Using [³H]-C₆ as a marker fortotal C₆, liposomal delivery resulted in a time-dependant accumulationof ceramide in caveolae lipid signaling rafts (FIG. 9B). Ceramideaccumulated in fractions No. 4 and 5 of a sucrose gradient, whichrepresented caveolin-1 enriched lipid rafts (caveolae). (FIG. 9B)

In Vivo Summary

Using an in vivo mouse model system of breast adenocarcinoma, a methodwas established for systemic delivery of C₆ for the treatment of solidtumors. In vivo data suggest promising anti-cancer activity withpegylated liposomal formulations in 410.4 tumor-bearing BALB/c mice(FIG. 10A-B). These in vivo results show a dose-responsive reduction intumor volume with liposomal C₆, compared to empty Ghost liposomes. Thisis the first study demonstrating any efficacy of systemic C₆formulations in an in vivo model of tumorigenesis. Moreover, in vivopharmacokinetic analysis demonstrated that systemic liposomal C₆delivery resulted in the attainment and maintenance of steady statebioactive concentrations of C₆ in tumor tissue over a 24 hour period(FIGS. 11A-B). This steady state bioactive concentration was maintainedalthough C₆ was rapidly cleared from the blood and major first-passorgans. Taken together, it is shown that systemic formulations of C₆displayed efficacy both in vitro and in vivo with favorablepharmacokinetics. Moreover, using Swiss Webster mice, pegylatedliposomal-C₆ formulations demonstrated no toxic side effects followingintravenous injection of up to 100 mg/kg, whereas the injection of freeC₆ in DMSO killed 50% of the mice at 10 mg/kg.

In Vivo Anticancer Efficacy

In order to assess the in vivo efficacy of systemic liposomal-C₆delivery, 5×10⁶ 410.4 cells were injected subcutaneously into the righthind flank of Balb/C mice. Four days following the injection of 410.4cells, mice were injected intravenously (i.v.) with either liposomal-C₆,empty liposomes (Ghost), or 0.9% NaCl. Mice were treated every two days.Immediately prior to treatment, mice were weighed and tumors weremeasured. Tumor size was measured with calipers and tumor volume wascalculated using the formula for a hemiellipsoid: V=π/6×L×W², whereV=tumor volume, L=length, and W=width.

In Vivo Pharmacokinetics

Liposomal-C6 [DSPC/DOPE/DSPC-PEG(5000)/C₈—PEG(750)/C₆-ceramide(3.75:1.75:0.75:0.75:3.0)] delivery displayed dose-dependent anti-tumoractivity via ceramide-induced apoptosis. Systemic delivery ofliposomal-C₆ inhibited tumor growth in a dose-dependent manner, comparedto empty ghost liposomes (FIG. 10A). Tumors were removed following oneweek of treatment with 40 mg/kg liposomal-C₆ and cryo-sections weregenerated for histological analysis (FIG. 10B). Tumor sections werestained with a TUNEL Kit to assess the degree of induced apoptosis;DAPI-stained nuclei.

In Vivo Pharmacokinetics

Using [H]-C₆ as a marker for C₆ delivery, tumor-bearing mice wereinjected with 10 and 40 mg/kg liposomal-C₆, and blood, tumors, spleen,kidney, liver, and heart tissue were removed at chosen timepoints.Tissues were weighed, solubilized, and counted using a scintillationcounter. The mass of total C₆ per mg of tissue (or ml of blood) wascalculated for each tissue taken and a pharmacokinetic profile wasevaluated. In order to trace the delivery of liposomal vehicles relativeto the distribution of C₆, [³H]-CHE was incorporated into liposomalformulations as a marker for the delivery vehicles.

Doses of 10 and 40 mg/kg liposomal-C₆ appeared to follow first orderkinetics, with a sufficient plasma concentration correlating to the invitro IC₅₀ sustained at 24 hours (FIG. 11A). At these doses, asteady-state concentration of C₆ in the tumor tissue was achieved atapproximately 30 minutes (FIG. 11B). The 40 mg/kg dose maintained aconcentration well above the desired IC₅₀ up to 24 hours. Using [³H]-CHEas a marker for the pegylated liposomal vehicle, the liposomes appearedto accumulated in the tumor tissue in a time dependent manner. This maysignify that the steady state concentrations of C₆ in tumor tissue maybe sustained due to continued accumulation of the pegylated liposomes,thus replenishing metabolized C₆ in the tumors.

EXAMPLE 2 Dendrimers as a C₆-Ceramide Drug Delivery Vehicle

Dendrimer Synthesis

Dendrimers were synthesized by conjugating poly(L-lysine) (PLL) dendronwith PNIPAAM grafted with PLLA. PNIPAAM grafted with PLLA wassynthesized by free radical polymerization. (FIG. 12).

Dendrimer Thermo-responsive Properties

UV-vis spectroscopy (Perkin Elmer Lamda 25, Shelton, Conn.) was used tostudy the transmittances of dendrimers at 500 nm in PBS (pH=7.4) withtemperature increase at 1° C./30 min at various concentrations (FIG.13A). The dendrimers were thermo-responsive, showing a lower criticalsolution temperature (LCST) (defined as temperature at 95% of maximumtransmittance) of 31, 32, 34, and 39° C. at concentrations of 1, 0.5,0.1, and 0.05 mg·ml⁻¹, respectively. The LCST became obscure withdecreasing concentration of the dendrimers. Above the LCST, thetransmittance magnitudes decreased with increasing concentration due tothe increase of the interactions of the polymers. The LCST of thePNIPAAM and the PNIPAAM grafted with PLLA decreased linearly withlogarithmic concentration, and the latter was 2° C. lower than theformer over the concentrations, due to the hydrophobicity of the PLLA.However, when PLL was conjugated at both ends of the PNIPAAM graftedwith PLLA, the LCST of the dendrimer showed a non-linear relationshipwith logarithmic concentration and the highest value compared to that ofother two types of polymers, due to the positive charges andhydrophilicity of the PLL.

The thermo-responsive properties of the dendrimers were confirmedfurther by measuring hydrodynamic sizes of the dendrimers againsttemperature using dynamic light scattering (DLS) (ALV, Germany). Theapparent hydrodynamic diameters (D_(h)) of the dendrimers in PBS(pH=7.4) at three concentrations 1, 0.5 and 0.1 mg·ml⁻¹ showed atemperature dependence in three regions, respectively (data not shown).In the lower temperature range, D_(h) decreased slightly as the solutiontemperature increased, reflecting the contraction of individual chains.In the middle temperature range, D_(h) increased before reaching theirmaximum values, showing that the dendrimer nanoparticles aggregated witheach other due to interchain association. In the higher temperaturerange, D_(h) decreased as the aggregation temperature increased due tointrachain contraction. The LCST of the dendrimers was 29, 30 and 31°C., defined as the initial break points of the D_(h)-temperature curves,at three concentrations: 1, 0.5 and 0.1 mg·ml⁻¹, respectively. In boththe lower and middle temperature ranges, D_(h) increased with increasingconcentrations because interchain interactions also increased withincreasing concentrations. The LCST determined by the DLS was slightlylower than that determined by the UV-vis spectroscopy for the samesolution concentration, attributed to different instruments for themeasurement. Both the DLS and UV-vis results demonstrated that the LCSTdecreased with increasing concentrations.

Dendrimer Degradation Properties

C₆ Ceramide and Bio-Smart Nanodendrimers

Dynamic degradation of dendrimers in PBS (pH=7.4) at 1 mg·ml⁻¹ at atemperature below and above the LCST, 25 and 37° C., respectively, wasprobed by measuring molar mass changes of the dendrimers as a functionof time using MALDI-TOF. The number molar mass (M_(n)) of the dendrimersdecreased with time for up to one month, and decreased faster at 37° C.than at 25° C., and reached a relatively stable value after 19 days atboth temperatures. Interestingly, the stable M_(n) after 19 days, wasaround 2700 g·mol⁻¹, and its subtraction from the initial M_(n) (around4200 g·mol⁻¹) was around 1500 g·mol⁻¹, which was equal to that of thePLLA. The results suggest that the dendrimers degraded, and theirdegradation might be attributed to the hydrolytical degradation of thePLLA component of the dendrimers. To support the above statement, theFTIR spectra and viscosity of the dendrimers as a function of time weremeasured, respectively (data not shown). It was observed that the peakintensities at ˜1760 cm⁻¹, which was due to the ester C═O stretching ofPLLA, clearly decreased with time and disappeared after 19 days. Becausethe peaks at ˜1660 cm⁻¹, which was attributed to amide C═O stretching ofPNIPAAM and PLL were relatively stable, they were used as referencepeaks to normalize the peak intensities at ˜1760 cm⁻¹. The resultantpeak height percentage decreased with time and became 0 after 19 days(data not shown). Additionally, it was observed that the viscosity ofthe dendrimer (measured by a Cannon-Ubbelohde type viscometer, followingthe procedures of ASTM D 445 and ISO 3104) decreased with time,decreased faster at 37° C. than at 25° C., and reached a stable valueafter 19 days (data not shown). Therefore, the FTIR results, togetherwith the viscometer and MALDI-TOF results, strongly suggested that thedesigned dendrimers were biodegradable due to the hydrolytic degradationof the PLLA component.

Methodology for C₆ loading efficiency. C₆ was mixed with the dendrimerat a ratio of 3:1 (C₆: dendrimer, w/w) in a solvent system comprised ofdistilled water, ethanol and N-dimethylformamide (DMF) (distilledwater/ethanol/DMF (5:5:3, v/v/v) at a concentration of 1 mg/ml, and wassealed and stored at room temperature for 7 h. The C₆/dendrimer solutionwas put into a cellulose membrane (MWCO-3500) and dialyzed againstethanol (50 ml) to remove free C₆ from inside the membrane. The amountof C₆ inside and outside the cellulose membrane was measured byMALDI-TOF mass spectrophotometry. Both solutions inside and outside themembrane were mixed with a matrix solution of 2,5-dihydroxybenzoic acidat 1:9 (sample:matrix). C₁₆-ceramide (C₁₆) was used as an internalstandard material and added to each solution. The amount of C₆, as afunction of time (2, 4, 6, and 10 h.), was calculated by the relativeintensity of C₆ and C₁₆ mass peaks at 424 and 562 m/z, respectively.Loading of C₆ into dendrimers at a ratio of 3:1 (C₆:dendrimer) resultedin a loading efficiency of approximately 35.9±1.2%.

Methodology for C₆ release from dendrimers. In order to assess theinteraction between C₆ and the dendrimer, it is necessary to evaluatethe release kinetics of C₆ from the dendrimer. The fractional release ofC₆ (M_(t)/W_(ω) where M_(t) and W_(ω) are the amount of the C₆ releasedat time t and the maximum amount of C₆ released, respectively) increasedwith time due to the hydrolytic degradation of the dendrimer. Thedendrimer-C₆ complex was dissolved in sterile PBS (pH=7.4) and put intoa cellulose membrane (MWCO=3500) and dialyzed against sterile PBS(pH=7.4) (50 ml) containing sodium dodecyl sulfate (SDS) at 0.5% (w/v).Since C₆ is extremely hydrophobic, it was essential to perform thedialysis in the presence of a detergent, such as SDS. Finalconcentrations of the dendrimer was 0.1 mg/ml with continuous magneticstirring at temperatures below (25° C.) and above the LCST (37° C.). Atselected time intervals (between 0 and 30 days), 1 ml buffer solutionwas removed and replaced with fresh buffer, in order to determine theconcentration of the released C₆. In order to quantitate C₆-release fromthe dendrimer, the amount of C₆ inside and outside the cellulosemembrane was measured by MALDI-TOF mass spectrophotometry, using C₁₆ asan internal standard. At 37° C., a temperature above the LCST of thedendrimer, the dendrimer is more hydrophobic, thus resulting in a slowerrelease profile of C₆ from the C₆-loaded dendrimer (FIG. 13B).

Dendrimers as a Drug Delivery Vehicle for C₆-ceramide

Dendrimer Degradation Properties

Design of Dendrimers with an LCST Above Physiologic Temperature

The primary objective is directed at generating biodegradable andtemperature-sensitive dendritic nanoparticles that can be complexed toC₆, are injected intravenously, are soluble in the blood stream for along period of time, and achieve targeted and sustained delivery oftherapeutic agents to solid tumors. In order to create dendrimers with aLCST of approximately 40° C. for thermally targeting the dendrimers tosolid tumors, the relative molar ratios between NIPAAM, hydrophobic PLLAand hydrophilic PLL play a critical role. It is well known that thehomopolymer PNIPAAM has a LCST of 32° C. {Eeckman, 2001 #52}. This LCSTwill decrease or increase by increasing the amount of the incorporatedhydrophobic or hydrophilic component, respectively{Eeckman, 2001 #52}.

The dendrimer structure was optimized in order to engineerbio-responsive, smart dendrimers that will release C₆ upon the inductionof local hyperthermia above physiological temperature. In order todesign a dendrimer with an optimal LCST slightly above 37° C., we havereplaced PLLA with a more flexible and amorphous polymer, such aspoly(D,L-lactic acid) (PDLLA) with molar masses 800, 2000 and 4000g/mol. Secondly, we used different generations (successive concentricrings of dendritic structure) of PLL, such as 3, 4 or 5. Finally, weused different molar ratios between PDLLA macromer and NIPAAM monomer,such as 0.02, 0.05 and 0.1 mol %. The LCSTs of the resulting dendrimerscan be assessed by UVvis spectroscopy by measuring transmittance as afunction of temperature and light scattering by measuring hydrodynamicsize as a function of temperature.

The LCST of the dendrimers was increased by copolymerization withN-isopropylmethacrylamide (NIMAAM). The resulting core-copolymersexhibited the LCST of 36, 42, and 44° C. with increasing NIMAAM monomerat 50, 60, and 70% of NIPAAM monomers, respectively. In this way, wehave designed bio-smart dendrimers that can be engineered to releasehydrophobic chemotherapeutic agents, including growth-arresting,pro-apoptotic lipid-derived second messengers to solid tumors throughtargeting the tumor via local hyperthermia. Localized heat may beapplied using ultrasound or heat patch devices to elevate the localtumor temperature above the LCST of the dendrimers. This process iscoined as “physiological hyperthermic drug delivery.”

Dendrimer Cell Viability

Due to the advantages that polymeric nanoparticles have compared toliposomal technologies as discussed above, we have loaded C₆ intotemperature-sensitive dendrimer nanoparticles in order to target solidtumors using a temperature-induced delivery strategy. Our proprietarydendrimer nanoparticles are comprised of PLLA, PLL, and PNIMPAM. Asstated above, using UV-vis spectroscopy to monitor the dendrimersolution transmittance with increasing temperature, we observed a sharptransition, confirming that these dendrimers are indeedthermoresponsive. The approximate LCST for these prototypic dendrimerswas found to be approximately 34° C. at 100 μg/ml. Using MALDI-TOF massspectrophotometry, we demonstrated that the molar mass of the dendrimersdecreased with time, verifying that they are also biodegradable. Asanalyzed by confocal microscopy, the dendrimers preferentiallyaccumulate into MDA cells at a temperature above the LCST (37° C.) thanbelow the LCST (25° C.) (FIG. 14A), likely due to increasedhydrophobicity. Moreover, using flow cytometry to quantitate theintracellular accumulation and uptake of FITC-labeled dendrimers, wedemonstrated that significantly more dendrimer uptake results in MDAcells at a temperature above the LCST (37° C.) than below the LCST (25°C.). (FIG. 14B). More importantly, treatment of MDA cells with theseC₆-loaded dendrimers resulted in the significant growthinhibition/cytotoxicity, while dendrimer alone displayed no cytotoxicity(FIG. 15A). In addition to inducing growth arrest, C6-enricheddendrimers induced apoptosis of MDA cells (FIG. 15B). Taken together,our preliminary data demonstrate that the dendrimers effectively controlrelease of an anti-cancer drug C₆. Optimization of this embodiment hasincluded the design of polymeric dendrimers with an LCST slightly abovephysiological temperature, to allow for physiological hyperthermic drugdelivery.

The following example is further illustrative.

EXAMPLE 3 Synthesis of Ceramide-containing Calcium Phospho-SilicateShell Resorbable Nanoparticles for Systemic Delivery

Reverse micelles were prepared using the nonionic surfactantpoly(oxyethylene) nonylphenyl ether (Igepal CO-520, Aldrich ChemicalCo.) without further purification. Cyclohexane, deionized water andC₆-ceramide was used as received for the synthesis.

Microemulsions of 20 mL total volume, consisting of 4 mL Igepal, 10 mLcyclohexane and deionized water, were prepared at ambient temperature ina 30 mL vial with rapid stirring. This produced a uniform mixture towhich trace amounts of C₆ was added in an aqueous phase as a micellularmixture of water and drug. The resulting micelle structure containing C₆was coated with CPS having composition Ca_(x)(PO₄)_(y) zSiO₂, where0.1≦x≦10 0.1, 0.1≦y≦10, and 0≦z≦10. The size of the resultingnanoparticles was controlled by varying the ratio of water to surfactant(R=[water]/[surfactant].

After the reverse micelles were encapsulated in the CPS coating, thenanoparticles were washed and concentrated using size exclusion highperformance liquid chromatography (SEC) modified specifically for theC₆-containing shell nanoparticles. An elution column shorter than normalwhich contains microporous silica particles having a diameter of about20 microns diameter was utilized in which ethanol is added as theelution solvent. To disperse the nanoparticles, an alkylamine silanecoupling agent, aminopropyltrichlorosilane, was added to the suspension.This coupling agent also was added to the microporous silica particlesused to pack the SEC column. The pH of the suspended nanoparticles wasmaintained at approximately 7.0 by adding acetic acid or sodiumhydroxide as needed. Additionally, a carbodiimide-mediated polyethyleneglycol (PEG) coupling agent was attached to the alkylamine couplingagent.

EXAMPLE 4 Hydrogels as C₆-Ceramide Release Vehicles

Nine multifunctional hydrogels with both thermoresponsive andbiodegradable properties were synthesized and characterized. Thehydrogels are copolymeric networks composed of N-isopropylacrylamide(NIPAAM) as a thermoresponsive component, poly(L-lactic acid) (PLLA) asa hydrolytically degradable and hydrophobic component, and dextran as anenzymatically degradable and hydrophilic component. Due to theirmultifunctional properties, the designed hydrogels are suitable forbiomedical applications including drug delivery and tissue engineering.(FIG. 16.)

The hydrogels showed thermoresponsive properties and the LCST was around32° C., typical to that of PNIPAAM. The hydrogels were alsohydrolytically biodegradable with pore sizes increasing after about 4months. The swelling behaviors of the hydrogels were different attemperature above (37° C.) and below (25° C.) the LCST and stronglydepended on the hydrophilicity and hydrophobicity of the copolymers. Inconclusion, the hydrogels have great potential for a controlled andsustained release of C₆-ceramide through changing their copolymercompositions and thermo-responsive and biodegradable properties.

It should be understood that the examples and embodiments describedherein are for illustrative purposes only and that various modificationsor changes in light thereof will be suggested to persons skilled in theart and are to be included within the spirit and purview of thisapplication.

The invention claimed is:
 1. A resorbable unagglomerated nanoparticledrug delivery vehicle for delivering a growth-arresting lipid-derivedbioactive compound and/or a hydrophobic chemotherapeutic agent and/orgene therapy agent to an animal or human in need of such delivery,comprising a dispersant attached to a calcium phospho-silicate shellencapsulating the lipid-derived compound, hydrophobic chemotherapeuticagent or gene therapy agent, wherein the lipid-derived compound isselected from the group consisting of C2-C10 ceramides and C18-C24ceramides, dimethyl sphingosine, trimethyl sphingosine, ether-linkeddiglycerides, ether-linked phosphatidic acids, sphingosines andsphinganines; and the gene therapy agent is selected from the groupconsisting of oligonucleotides, ribozymes, DNA-zymes, plasmids,antisense or Si-RNA therein.
 2. The resorbable unagglomeratednanoparticle drug delivery vehicle of claim 1, wherein the dispersant isselected from the group consisting of citric acid, tartaric acid, aceticacid, aminopropyltrichlorosilane, 3-aminopropyltrimethoxysilane (APS),3-aminopropylsilsesquioxane, 3-glycidoxypropyl-trimethoxysilane (GPS),trimethoxysilyl-propyldiethylenetriamine (DETA),3-trimethoxysilypropylsuccinic anhydride, and alkylcarboxylic acidsilane coupling agents.
 3. The resorbable unagglomerated nanoparticledrug delivery vehicle of claim 2, wherein a carbodiimide-mediatedpolyethylene glycol (PEG) coupling agent is added to the dispersant inorder to provide an attachment point for binders.
 4. A method ofadministering a composition to an animal or human in need thereof,comprising administering a composition comprising the resorbableunagglomerated nanoparticle drug delivery vehicle of claim 1 to theanimal or human, wherein the growth-arresting lipid-derived bioactivecompound or gene therapy agent contained therein is in a concentrationof up to approximately 1000 mg ml⁻¹.
 5. The method of claim 4 whereinthe growth-arresting lipid-derived bioactive compound and/or ahydrophobic chemotherapeutic agent and/or gene therapy agent comprisesat least approximately 40 wt % of the composition.
 6. A resorbableunagglomerated nanoparticle for delivering a growth-arrestinglipid-derived bioactive compound and/or a hydrophobic chemotherapeuticagent and/or gene therapy agent to an animal or human in need of suchdelivery, comprising a dispersant attached to a calcium phospho-silicateshell encapsulating the lipid-derived compound, hydrophobicchemotherapeutic agent or gene therapy agent, wherein the lipid-derivedcompound is selected from the group consisting of C2-C24 ceramides,dimethyl sphingosine, trimethyl sphingosine, ether-linked diglycerides,ether-linked phosphatidic acids, sphingosines and sphinganines and thegene therapy agent is selected from the group consisting ofoligonucleotides, ribozymes, DNA-zymes, plasmids, antisense or Si-RNAtherein.
 7. The resorbable unagglomerated nanoparticle of claim 6 havinga diameter ranging from between about 1 to 300 nm.
 8. The resorbableunagglomerated nanoparticle drug delivery vehicle of claim 7, whereinthe resorbable unagglomerated nanoparticles have a diameter of about 20nm or less.
 9. The resorbable unagglomerated nanoparticle of claim 6having a diameter less than 50 nm.
 10. The resorbable unagglomeratednanoparticle of claim 9 having a diameter of about 20 nm or less. 11.The resorbable unagglomerated nanoparticle of claim 6, wherein thedispersant is selected from the group consisting of citric acid,tartaric acid, acetic acid, aminopropyltrichlorosilane,3-aminopropyltrimethoxysilane (APS), 3-aminopropylsilsesquioxane,3-glycidoxypropyl-trimethoxysilane (GPS),trimethoxysilyl-propyldiethylenetriamine (DETA),3-trimethoxysilypropylsuccinic anhydride, and alkylcarboxylic acidsilane coupling agents.
 12. The resorbable unagglomerated nanoparticleof claim 11, wherein a carbodiimide-mediated polyethylene glycol (PEG)coupling agent is added to the dispersant in order to provide anattachment point for binders.
 13. The resorbable unagglomeratednanoparticle of claim 11 wherein the alkylcarboxylic acid silanecoupling agents are amide linked carboxyl groups.
 14. The resorbableunagglomerated nanoparticle of claim 12 wherein the binders areantibodies.